Detector systems for integrated radiation imaging

ABSTRACT

Detector systems for enhanced radiographic imaging incorporate Compton and PET imaging capabilities. The detector designs employ one or more layers of detector modules comprised of edge-on or face-on detectors, or a combination of edge-on and face-on detectors, which may employ structured detectors. The detectors implement tracking capabilities and operate in a non-coincidence or coincidence detection mode.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.13/573,981, filed Oct. 18, 2012, which claims priority to U.S.Provisional Application Ser. No. 61/689,139, filed May 31, 2012, andU.S. Provisional Application Ser. No. 61/690,348, filed Jun. 25, 2012,each of which is incorporated by reference herein, in the entirety andfor all purposes.

FIELD

This invention provides novel implementations of Compton camera detectorsystems that can be employed as novel integrated Compton-PET detectorsystems and CT-Compton-PET detector systems for diagnostic medical andsmall animal imaging (x-ray CT, nuclear medicine, PET) as well asradiation therapy, dental, industrial, Homeland Security and scientificradiation imaging.

BACKGROUND

The combining of imaging modalities to offer increased functionality hasproduced a number of useful imaging systems, particularly in medicaldiagnostic and small animal imaging. For example, Gamma ray PET detectorsystems are frequently sold with x-ray CT detector systems (although thePET and CT detector systems are physically separate and therefore do notshare detectors or a common imaging space). A notable previous attemptat offering an integrated imaging system (in which detectors and theimaging space of the system are shared) was a SPECT-PET (nuclearmedicine and PET) imaging system which reduced costs by sharingdetectors and the imaging space (the volume in which the object isimaged).

Although these SPECT-PET imaging systems were not well receivedcommercially due to performance compromises nonetheless they offeredinteresting functionality since SPECT and PET images could be acquiredseparately or simultaneously in a shared imaging space (thereby avoidingregistration error between separately acquired SPECT and PET images andreducing the total scan time). In addition, simultaneous CT-SPECTsystems have been proposed (typically using CZT or CdTe) although issuesarise due to generally differing collimation and flux rate requirements.

SUMMARY

The invention utilizes the recent improvements in high speed detectorelectronics along with detector materials developed for medicaldiagnostic slit scanning and CT, nuclear medicine, PET imaging, highenergy physics, inspection, etc. to develop cost-effective, singlepurpose and multipurpose integrated detector systems which implement oneor more properties of Compton camera detector systems. Compton camerasare frequently implemented as multilayer detectors. Conventionalphoton-tracking Compton camera designs include a single layer (afront-end detector) which provides 3D detector properties byincorporating a stack of face-on detector planes of the same materialsuch as low-Z Silicon (Si) or moderate-Z Germanium (Ge), essentially amultilayer detector, and a multilayer (dual layer) configuration whichcombines a 2D detector first layer (the front-end detector) and a 2Ddetector second layer (the back-end detector).

The conventional dual-layer, front-end/back-end detector configurationtypically consists of a face-on, planar, 2D Si (low-Z) front-enddetector combined with a face-on, 2D high-Z back-end detector. Thus,these two Compton camera configurations utilize detector layers of thesame material (low-Z and moderate-Z for Compton scattering) or differentmaterials (low-Z for Compton scattering and high-Z for photoelectricinteractions) for the detection of photons in the diagnostic energyrange of medical imaging.

Clearly other choices of materials can be made depending on the photonenergy range or if other types of particles (neutrons, muons, etc.) areto be detected. Unconventional Compton camera designs (as well as x-rayscanning and CT, SPECT, PET and hand-held probe designs have beendescribed in various U.S. patents and applications including: Nelson,U.S. Pat. No. 4,560,882; Nelson, U.S. Pat. No. 4,937,453; Nelson, U.S.Pat. No. 5,258,145; Nelson, U.S. Pat. No. 6,583,420; Nelson, U.S. Pat.No. 7,291,841; Nelson, U.S. Pat. No. 7,635,848; Nelson, U.S. Pat. No.8,017,906; Nelson, U.S. Pat. No. 8,115,174; Nelson, U.S. Pat. No.8,115,175; Nelson, U.S. Pat. No. 8,183,533; Nelson, U.S. patentapplication Ser. No. 13/199,612, U.S. Publication No. 2012/0181437;Nelson, U.S. patent application Ser. No. 13/507,659, U.S. PublicationNo. 2013/0028379; and are incorporated by reference herein.

Compton camera detector systems exploit the Compton scatter interactionand can also exploit photoelectric interactions (and even pairproduction interactions at sufficiently high photon energies). Comptoncamera detector systems include the capability to track theseinteractions in terms of spatial location and energy deposition with atemporal resolution limited by the detector itself and the readoutelectronics. Typically the interaction information is used to estimatethe directionality and energy of the photon incident on the Comptoncamera detector system whether the photon is an x-ray, a gamma ray, oran annihilation gamma ray.

Note that with the addition of collimation such as (for example) apinhole or parallel hole collimator, the Compton camera can be convertedinto a nuclear medicine SPECT camera (Gamma camera). This is an exampleof a dual-use, integrated Compton detector system in which the types ofapplications are relatively unchanged but the capabilities of the of thedetector system are modified (Nelson, U.S. Pat. No. 7,291,841; Nelson,U.S. Pat. No. 8,017,906). The collimation now provides thedirectionality of an incident gamma ray independent of directionalitydetermined by applying Compton camera reconstruction algorithms. It willbe shown that the integrated Compton detector system design can beapplied to a range of applications (including nuclear medicine).

By employing two or more Compton camera detector systems with electroniccoincidence circuitry (used in medical PET detector systems) coincidencePET imaging can be implemented. The flexibility of the Compton cameradetector system design allows versatile non-coincidence Compton-PET andcoincidence Compton-PET detector systems to be implemented. Furthermore,CT capability can be implemented in the Compton camera detector systemdesign, including non-coincidence and coincidence Compton-PET designsresulting in CT-Compton-PET detector systems. A simplification of thisdesign in which the CT detector and the Compton-PET detector (or just aPET detector) function independently will be referred to as a limitedCT-Compton-PET detector system.

Although applications discussed herein are primarily directed at medicaldiagnostic x-ray and gamma ray radiation detection, in principle theinvention can also be used to detect radiation such as charged particles(alphas, betas, protons, muons, etc.) and neutrons (as well as otherneutral particles) for the applications previously described.Furthermore, the Compton camera detector systems described herein can becombined with or integrated with other imaging modalities such as MRIscanners, optical scanners, ultrasound scanners, opto-acoustic scanners,microwave scanners, etc. It should be understood that the variations ofthe dual-use detector systems (triple-use detector systems can alsoimplemented) described herein can be employed for simultaneous ornon-simultaneous imaging as required by the appropriate application.

The invention provides Compton camera detector designs that employ oneor more layers of detector modules comprised of edge-on or face-on (ortilted) detectors or a combination of edge-on and face-on detectors (aswell as tilted detectors). Edge-on detectors (and tilted detectors) canincorporate sub-aperture resolution (SAR) capabilities and face-ondetectors can incorporate depth-of-interaction (DOI) capabilities. Oneor more types of detectors can be employed, including: scintillatordetectors, semiconductor detectors, gas detectors, low temperature (suchas Ge or superconductor) detectors and structured detectors. Detectorscan offer block, 1D, 2D or 3D spatial resolution as well as adequate,fast or very fast temporal resolution (depending on the applicationrequirements).

Detectors can offer fixed or adjustable pixels sizes which can beuniform or non-uniform. The effective pixel length along a detectorcolumn can be synthesized from the outputs of one or more uniformlyspaced pixels. Parallel or focused pixel structures can be implemented.Detectors can operate as energy integrators, photon counters (PCs) andphoton counters with energy resolution (PCEs). Possible detector formatsinclude planar (and focused planar) and focused structure (ring, partialring as well as focused ring and focused partial ring) detectorgeometries.

The details of one or more embodiments of the invention are set forth inthe accompanying drawings and the description below. Other features,objects, and advantages of the invention will be apparent from thedescription and drawings, and from the claims.

All publications, patents and patent applications cited herein arehereby expressly incorporated by reference for all purposes.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates a perspective view of a non-coincidence Compton-PETdetector imaging system.

FIG. 2 illustrates a perspective view of an edge-on silicon detectorsubstrate in which shielded readout ASICs are mounted within an etchedregion along the bottom edge of the semiconductor detector substrate.

FIG. 3 illustrates a perspective view of a focused planar detector.

FIG. 4 illustrates a perspective view of a coincidence Compton-PETdetector imaging system.

FIG. 5 illustrates a perspective view of a non-coincidenceCT-Compton-PET detector imaging system.

FIG. 6 illustrates a perspective view of a minifying scintillating fiberarray coupled to a 1D photodetector structured detector suitable for PCor limited PCE CT imaging.

DETAILED DESCRIPTION

The invention provides novel Compton camera detector designs and systemsfor enhanced radiation imaging including Compton and nuclear medicineimaging, PET imaging and x-ray CT imaging. The invention also providesintegrated detector systems based on Compton camera designs. In oneaspect, the invention provides integrated non-coincidence Compton-PETdetector imaging systems.

In another aspect, the invention provides integrated coincidenceCompton-PET detector imaging systems. In yet another aspect, theinvention provides limited integrated CT-Compton-PET detector imagingsystems. In still another aspect, the invention provides integratednon-coincidence CT-Compton-PET detector imaging systems. In anotheraspect, the invention provides integrated coincidence CT-Compton-PETdetector imaging systems. Since the integrated nature of these Comptoncamera detector design implementations is readily apparent the term“integrated” will frequently be omitted when referring to them.Therefore ‘integrated non-coincidence Compton-PET detector imagingsystems’ will also be referred to as ‘non-coincidence Compton-PETdetector imaging systems’, etc.

The invention employs a range of detector types and formats. The use ofgas, scintillator, semiconductor, low temperature (such as Ge andsuperconductor) and structured detectors in edge-on and/or face-ongeometries has previously been described for both medical andnon-medical imaging applications. Medical imaging applications includediagnostic x-ray imaging (such as slit scanning, slot scanning, flatpanel or planar cone beam CT, focused structure ring or partial ring fanbeam CT), nuclear medicine imaging (Compton camera, SPECT/gamma cameradetector imaging systems as well as hand held probe detectors), PETimaging, dental imaging and radiation therapy portal and cone beam CTimaging. Non-medical imaging applications include high energy Physics,x-ray and gamma ray Astronomy, industrial radiography, Homeland Security(HLS) and military applications.

Furthermore it has been shown that detector spatial resolution can beenhanced using sub-aperture resolution (SAR) or depth-of-interaction(DOI) readout techniques with edge-on and face-on detector geometries,respectively. Detectors may be layered (stacked) and detector moduleswithin a layer can be partially- or completely-offset from neighboringdetector modules. Detectors may function as energy integrators, photoncounters (PCs) or photon counters with energy resolution (PCEs)depending on the application.

High speed electronics is provided for tracking interactions andanalyzing the readout signals. An electronic communications link isprovided to a computer for data post-processing, storage, and display.

One or more tracking capabilities such as examining nearest neighborpixels for effects related to induced signals and charge diffusion,following scattered or characteristic x-ray radiation within a detectorlayer and between detector layers (if there is more than one detectorlayer), following Compton scattered electrons and photoelectrons andmeasuring coincidence events (for example, the detection of pairs ofannihilations photons in PET imaging), etc. can be implemented. Trackingtechniques are used in photon counting and spectral x-ray imaging,SPECT, PET, Compton cameras, hand-held radiation detector probes,neutron detectors, detectors with SAR or DOI capability and high energyPhysics particle detectors.

Various Compton camera implementations have been described previouslythat incorporate one or multiple detector layers. These detector layersprovide suitable 2D or 3D spatial resolution, energy resolution,temporal resolution and tracking capability. Compton camera, nuclearmedicine SPECT/gamma camera and PET detector imaging systems, tracking,x-ray CT and slit and slot scan detectors, hand held probe detectors,edge-on and face-on detectors (with or without SAR or DOI capability),integrating, PC, and PCE detectors, multi-material detectors along withplanar and focused structure detector geometries have been described invarious U.S. patents and applications including: Nelson, U.S. Pat. No.4,560,882; Nelson, U.S. Pat. No. 4,937,453; Nelson, U.S. Pat. No.5,258,145; Nelson, U.S. Pat. No. 6,583,420; Nelson, U.S. Pat. No.7,291,841; Nelson, U.S. Pat. No. 7,635,848; Nelson, U.S. Pat. No.8,017,906; Nelson, U.S. Pat. No. 8,115,174; Nelson, U.S. Pat. No.8,115,175; Nelson, U.S. Pat. No. 8,183,533; Nelson, U.S. patentapplication Ser. No. 13/13/199,612, U.S. Publication No. 2012/0181437;Nelson, U.S. patent application Ser. No. 13/507,659, U.S. PublicationNo. 2013/0028379; and are incorporated by reference herein.

X-ray or gamma ray interactions (in medical imaging applications) can betracked between sufficiently thin detector layers, each with 2D spatialresolution capability. If the depth of a 2D detector layer issufficiently small such that tracking position errors are acceptablethen it effectively functions as a restricted 3D detector (its depthresolution is at most the thickness of the detector layer).

If detectors offer 3D spatial resolution capability then interactiontracking (including multiple interactions) can be implemented internallywithin a 3D detector layer as well as between detector layers. Energyresolution can be used to measure the position-dependent energy lossesdue to the interactions within detectors which in turn can provide anestimate of the energy of the initial incident x-ray or gamma ray. Thisinformation can be used to determine whether the initial incident photonenergy is within an allowed energy range as well as its directionality.Temporal resolution capability can be used to distinguish betweenindependent incident x-rays or gamma rays interactions (as well as theirsubsequent interactions) within the Compton camera.

It will be shown that very good temporal resolution can be beneficial ifcoincidence timing is of interest between Compton camera systems (forexample, when coincidence PET imaging is implemented). One well-knownimplementation of a Compton camera reduced cost by using a dual-layerdetector design wherein the first layer (front-end) was a small area,face-on, Si or Ge semiconductor pixelated detector which offered 2Dspatial resolution. The second layer (back-end) was a large area,face-on, scintillator (gamma camera) detector which also offered 2Dspatial resolution (Singh, M., Medical Physics Vol. 10(4), pp. 421-427,July/August 1983 and Singh, M., Doria D., Medical Physics Vol. 10(4),pp. 428-435, July/August 1983).

Both front-end and back-end detectors offered appropriate levels ofenergy resolution for the photon energies employed and temporalresolution for the expected event interaction rates. Since Comptonscattered photons include a range of scatter angles the sensitivity ofthis design is in-part dependent on the separation distance and area ofthe second layer from the first layer of detectors. A second layer whichemploys a smaller 3D detector may, in some instances, be more-costeffective than a larger 2D detector which suffers from parallax errorsand needs to be positioned further away from the first layer.

Another implementation of the Compton camera, the conventional (face-on)Compton telescope camera, consisted of only a first layer detector. Thisfront-end detector was comprised of a stack (and thus could also beviewed as a multilayer detector) of 2D, face-on Si detectors whichfunction together as a 3D detector (Kroeger R, et al., IEEE Trans. Nucl.Sci., Vol. 49(4), pp. 1887-1892, August 2002; Nelson, U.S. Pat. No.8,017,906).

A stack of 2D, face-on Ge detectors (or a thick 3D Ge detector with DOIcapability) can also be implemented although the Ge semiconductor mustoperate at a low temperature. The Compton telescope camera tracksmultiple Compton scatters by a photon in order to determine its originaldirection and energy. Low Z (atomic number) semiconductor materials suchas Si and diamond (and sometimes moderate-Z Ge) are often preferred forthe front-end Compton scatter detector for photons of relatively lowenergies (e.g. medical diagnostic x-ray energies, 140.5 keV gamma raysfrom Tc-99m used in nuclear medicine) compared to the 511 keV gamma raysused in PET imaging.

The Compton scatter interaction cross section of the material dominatesits photoelectric cross section and the relative contribution to theangular reconstruction error due to the Doppler shift is reduced as Zdecreases and/or photon energy increases. As the photon energy increasessemiconductor materials with moderate-Z values (such as Ge, GaAs, CdTe,CZT, etc.) represent increasingly acceptable substitutes for low Zsemiconductor materials such as Silicon. The amount of energy depositedby relatively low energy photons (commonly used in diagnostic x-rayimaging or nuclear medicine) due to a Compton scatter interaction istypically small and therefore semiconductors detectors are employed asfront-end detectors because of their superior energy resolution comparedto most scintillator detectors.

In the dual-layer Compton camera design lower-cost 2D scintillatordetectors may be employed in place of semiconductor detectors asback-end detectors if they offer suitable spatial, temporal and energyresolution. The semiconductor front-end detector may be replaced by ascintillator (or gas) front-end detector although energy resolution maysuffer. Any significant reduction in accuracy of the calculated incidentphoton directionality by Compton reconstruction algorithms can beaugmented or supplanted by additional information such as coincidencebetween detectors (used in coincidence PET imaging).

Compton electron tracking in a gas detector can be implemented althoughthis is typically very time-consuming. Cherenkov radiation, despite therelatively weak optical signals, can be exploited for time-of-flight(TOF) measurements. (Cherenkov radiation can be detected when generatedin optically-transparent mediums including fluids such as liquids andgases, scintillators and non-scintillators such as transparent plastics,glasses, fibers, diamond films, etc. Thus, transparent dielectricmediums other than scintillators and gases can be also be employed asCompton scatter or photoelectron detectors within a Compton cameradetector system although energy resolution would suffer based on thedetection of Cherenkov radiation alone. Inexpensive dielectric materialsmay be acceptable for those applications in which radiation scatterwithin the object is of reduced importance and therefor lower detectorenergy resolution is acceptable. Variations of detector designsdescribed previously and herein can include measuring only a Cherenkovsignal or a Cherenkov signal and a fluorescence signal or an electronicsignal.)

Potential advantages of this dual-layer design may include aless-expensive front-end detector and/or a front-end detector thatoffers a feature such as fast (greater than 1 nanosecond) or very fast(less than 1 nanosecond) temporal resolution. Very fast temporalresolution is of interest for TOF PET. Benefits of TOF PET includeimproved image resolution and lower patient radiation dose. Furthermore,the use of coincidence information can also simplify the requirements ofthe back-end detector.

Compton electron tracking can also be implemented within a detectorlayer and between detector layers that employ at least one ofscintillator-photodetector, semiconductor, structured and lowtemperature detectors. Since electrons readily interact with matterelectron tracking is preferably implemented when detecting energeticphotons which are Compton scattered, typically generating more-energeticelectrons with a more-directional nature. (A similar concept applies toenergetic photoelectric interactions which typically generatemore-energetic photoelectrons with a more-directional nature. Thus, aCompton camera could utilize sufficiently energetic photoelectricinteractions for image reconstruction by tracking the highly directionalphotoelectrons.)

The tracking of Compton scattered electrons as well as Compton scatteredphotons can be simplified by enabling longer path lengths for thescattered particles, improving the estimates of scattering angles.Examples of relatively thin, edge-on detector configurations thatincorporate gaps between adjacent detectors (including partially- orcompletely-offset detectors) are shown in FIGS. 1, 3, 5. Face-ondetector configurations with gaps between detector layers can also beimplemented. Compton camera image reconstruction can be improved if boththe Compton scattered photon and electron are both tracked since thesolution can be limited to a fraction of a cone surface rather than thefull cone surface.

The flexibility of the Compton camera design can be understood byconsidering front end (single layer) detector and front-end withback-end (dual-layer) detector implementations of multilayer, edge-ondetector Compton camera designs which can be used for low energy andhigh energy photon imaging. In one dual-layer implementation thefront-end detector is used to detect low energy x-rays or gammas and theback-end detector acts as an edge-on SPECT/gamma camera or PET camera(Nelson, U.S. Pat. No. 7,291,841).

A focused, edge-on Compton camera design was described that can employone or multiple (of the same or different materials) detector layers aswell as implementing additional features such as the offset (complete orpartial) of detector modules within a layer. Completely offset detectormodules can be used to create two or more detector layers (offsetlayers) which when employed together can approximate a continuousdetector. The offset layer feature of an edge-on Compton camera designcan be implemented in PC, PCE and energy integration versions ofdiagnostic CT detector (as described in Nelson, U.S. Pat. No. 7,291,841;Nelson, U.S. Pat. No. 7,635,848; Nelson, U.S. Pat. No. 8,017,906;Nelson, U.S. Pat. No. 8,115,174; Nelson, U.S. Pat. No. 8,115,175;Nelson, U.S. Pat. No. 8,183,533; Danielsson, U.S. Patent Publication No.2010/0204942; Bornefalk, U.S. Patent Publication No. 2010/0215230).

Implementations of the Compton camera design are described herein thatare suitable for use as Compton-PET imaging systems and CT-Compton-PETimaging systems. In addition, the positioning of nuclear medicinecollimator hardware such as focused, parallel or pinhole collimatorsbetween the object being imaged and the Compton camera permits thesystem of collimator and Compton camera to provide conventional nuclearmedicine imaging capabilities (the imaging capabilities of a SPECT/Gammacamera) for those applications in which the Compton camera does notoffer adequate imaging properties.

Limited implementations of the Compton camera designs described hereininclude versions that function only as CT or PET (and SPECT) detectordesigns. The Compton camera imaging systems described herein will finduse in diagnostic medical x-ray CT, nuclear medicine and PET imaging,x-ray micro-CT imaging, dental CT, medium and small animal imaging,radiation therapy imaging, industrial imaging, HLS and military imagingand scientific imaging.

Compton-PET Detector Systems.

One implementation of the Compton camera is referred to as theCompton-PET detector system (Nelson, U.S. Pat. No. 7,291,841). TheCompton-PET detector system design allows flexibility in the choice ofdetector materials as well as detector geometries. This flexibility isconstrained by the intended imaging applications (such as PET only,nuclear medicine and PET, x-ray and PET).

Face-on, edge-on, and combinations of face-on and edge-on detectors canbe employed. One, two or more than two layers of detectors can beemployed. Detector modules that comprise a detector layer can optionallybe partially-offset or completely-offset from their neighbors within alayer. Common PET image acquisition formats based on planar and focusedstructure (such as ring and or partial ring) geometries are readilyimplemented.

Compton-PET detector systems are based on block, 1D, 2D or 3D edge-on,face-on, or mixtures of edge-on and face-on detectors (including edge-ondetectors with SAR capability and face-on detectors with DOI capability)previously described (Nelson, U.S. Pat. No. 4,560,882; Nelson, U.S. Pat.No. 4,937,453; Nelson, U.S. Pat. No. 5,258,145; Nelson, U.S. Pat. No.6,583,420; Nelson, U.S. Pat. No. 7,291,841; Nelson, U.S. Pat. No.7,635,848; Nelson, U.S. Pat. No. 8,017,906; Nelson, U.S. Pat. No.8,115,174; Nelson, U.S. Pat. No. 8,115,175; Nelson, U.S. Pat. No.8,183,533). The non-coincidence and coincidence Compton-PET detectorsystems described herein include focused and unfocused planar detectorformats and focused structure (such as ring and partial ring as well asfocused ring and focused partial ring) detector formats.

A non-coincidence Compton-PET (one-sided PET) detector system isimplemented by extending Compton camera designs that have previouslybeen developed for nuclear medicine imaging devices such as hand heldprobes or SPECT/Gamma cameras so that the detector system can operatewith the lower gamma ray energies used in nuclear medicine as well asthe higher energy range of PET with good detection efficiency.

A highly flexible implementation of a Compton camera design is adual-layer, 3D Compton camera. A specific implementation, anon-coincidence Compton-PET detector system, employs a (preferably, butnot exclusively) Compton scattering front-end detector and a(preferably, but not exclusively) high-stopping power back-end detectorin which both front-end and back-end detectors offer suitable 3D spatialresolution, energy resolution and temporal resolution (Nelson, U.S. Pat.No. 8,017,906).

Both the front-end and back-end 3D detectors provide adequate temporalresolution for an expected event rate such that accurate event trackingcan be enabled both within the front- and back-end detectors and betweenthe front-end and back-end detectors. Both the front-end and back-end 3Ddetectors can record Compton scatter and photoelectric interactions. Insome instances Raleigh scattering interactions (angle change withinsignificant energy loss) can be identified based on trackinginformation.

The front-end and back-end detectors, either separately or together, canoperate as two layer Compton cameras and Compton telescope cameras(Nelson, et al., U.S. Pat. No. 8,017,906). In one scenario the 3Dfront-end detector can function as a single (or multiple) Comptonscatter device and the 3D back-end detector can be used to measure theenergy and interaction location of the Compton scattered photon.

The front-end and back-end detectors have 3D spatial resolution.Front-end and back-end 3D detectors can also Compton-scatter a photon(measuring position and energy deposited) and detect the (single ormultiple) Compton-scattered photon (measuring its energy and interactionlocation). Therefore this two layer Compton camera with 3D detectorlayers incorporates the capabilities of three conventional two layerCompton cameras (in which one layer Compton-scatters the photon and theother layer detects (stops) the Compton-scattered photon).

Compton telescope camera designs exploit multiple Compton scattering forreconstruction. The Compton telescope camera capability can beimplemented in the 3D front-end detector, in the 3D back-end detectorand between the 3D front-end and back-end detectors (providing thecapabilities of three conventional (multilayer, face-on 2D arraydetectors) Compton telescope cameras. Appropriate two layer Comptoncamera and Compton telescope camera reconstruction algorithms are usedto form an image.

When this Compton camera is used to image single annihilation gamma rayscreated during a PET scan it is referred to as a one-sided PET detectorsystem or a non-coincidence Compton-Pet detector system. (Thisdual-layer, 3D Compton camera design is clearly not limited to PETimaging alone and therefore may be adapted for use in imagingapplications at other photon energies. Furthermore more than two layersof 3D detectors can be employed and non-3D layers of detectors can bemixed with 3D layers of detectors, thereby introducing additionalflexibility in the types of imaging applications for which this Comptoncameras design is suitable.)

This one-sided PET detector can be implemented in a focused or unfocusedplanar detector geometry or a focused structure detector geometry suchas a ring or partial ring (as well as focused ring and focused partialring detector geometry). This avoids the expense of employing acoincidence PET detector system based on opposing (or nearly-opposing)sets of PET detectors.

FIG. 1 shows a dual-layer Compton-PET detector imaging system 1000 thatincorporates 3D, edge-on detector arrays 510 and 520 (a first layer ofdetectors and a second layer of detectors, respectively). Theindividual, 2D edge-on detector modules 102 use crossed strip radiationdetectors 115. Alternatives include 2D pixelated arrays (or 3D pixelatedarrays if SAR capability is enabled) in an edge-on geometry.

Incident radiation photons 107 from gamma ray radiation source (notshown), with an energy less than the pair production threshold, canundergo Rayleigh scattering, Compton scattering or photoelectricinteractions. Compton scattered gamma ray photons 108 can be detected bythe edge-on radiation detector within the module 102 responsible for theinitial scattering, by other edge-on detectors modules within thefront-end detector layer 510 (detector layer 1) or by detector moduleswithin the back-end detector layer 520 (detector layer 2).

Each module 102 also includes a base 106 and a communications link 103.The base 106 preferably contains detector electronics including signalconditioners and readout ASICs, power management components, temperaturecontrol components, and a data or information channel for communicatingwith the computer system. The communications link 103 can be used toprovide power to the module 102 and connects the base 106 to a computersystem. The communication link 103 preferably is used to off-load thedigitized detector radiation data to a computer system for analysis andimage reconstruction.

The computer system, which can include general purpose, dedicated, andembedded computers, provides monitor and control services to modules102, to the detector layers 510 and 520 and to the entire Compton-PETdetector imaging system 1000. The computer system evaluates moduleparameters, detector layer parameters, and the detected radiation imagedata. The detected data is processed and can be displayed and stored ifdesired.

Additional relevant module information, such as temperature, amplifiersettings, detector voltages, position, orientation, and motioninformation, can be transmitted to this computer system over thecommunication link 103. The computer system transmits instructions thatupdate the detector modules 102 and detector layers 510 and 520. Thisestablishes a dynamic information feedback loop that is useful foradaptive imaging (Nelson, et al., U.S. Pat. No. 7,291,841).

Note that the electronic functionality of the detector base 106 can beimplemented along the side of a detector module or attached to thesurface of the detector module (integrated electronics). Another optionwhen the detector substrate is a semiconductor such as Si is to etch anindentation along the bottom of (opposite the radiation entrancesurface) and mount the readout ASICs and radiation shielding in theindentation and directly to the substrate along the bottom edge.

If the length of the edge-on detector is greater than its height thenthis configuration allows the readout ASICs to be closer to a set ofdetector pixels than for the case wherein the readout ASICs are mountedalong the side in order to avoid the direct x-ray beam. Preferably thecombined thickness of the etched substrate and mounted readout ASIC withshielding would be less than or equal to the thickness of the substrate(avoiding problems if the readout ASIC and any shielding stick out fromthe substrate and possibly interfering with the x-ray beam seen byoffset detectors).

FIG. 2 shows a perspective view of readout ASICs 200 with radiationshielding 204 mounted in an etched Si substrate 208 (or another suitablesemiconductor substrate) with a pixel size 215 that varies with heightwhich is positioned edge-on to incident radiation photons 109. Othermeans for delivering power to the detector modules as well as wirelesscommunication can be employed in place of communication link 103 (FIG.1).

It should be understood that readout ASICs can be mounted along the sideand the bottom edge. Two or more non-coincidence Compton-PET detectorsystems (an enhanced non-coincidence Compton-PET detector system) can beemployed for a PET imaging application. Furthermore, with the additionof coincidence circuitry, pairs of non-coincidence Compton-PET detectorsystems (preferably facing each other and positioned on opposite sidesof an object) can function as a coincidence Compton-PET detector system.

The cost of a two layer non-coincidence Compton-PET (one-sided PET)detector system can be reduced if either one or both of the 3D front-endand back-end detectors can be replaced by a suitable 2D detector withacceptable energy and temporal resolution. The caveat is that photondetection efficiency and reconstruction image quality may suffer as aresult. A compromise in terms of cost is to leave the front-end detectorwith 3D spatial resolution (and therefore retaining the previouslylisted capabilities: to function as a Compton scatterer, a two layerCompton camera, a Compton telescope camera) and employ a back-enddetector with 2D spatial resolution. The back-end detector would offeracceptable stopping power, energy resolution and temporal resolution forthe expected gamma ray event rate and gamma ray energies.

For a planar detector geometry the front-end and back-end detectors canconsist of single-layer face-on detector plane modules, a multilayer(stack) of face-on detector plane modules, a single-layer of edge-ondetector modules, a stack of edge-on detector modules or a combinationof face-on and edge-on detector modules. Face-on detector modules caninclude DOI capability whereas edge-on detector modules can include SARcapability.

One implementation of a focused planar detector geometry employs afront-end detector that consists of either a single layer (offset ornon-offset) or multiple layers (offset or non-offset) of tilted edge-ondetector modules. As an alternative to a parallel pixel structure afocused pixel structure can be implemented along the lengths of theedge-on tilted (or parallel) detector modules to account for x-ray beamdivergence.

FIG. 3 shows a perspective view of a focused planar detector system 1000in which detector modules 102 are tilted so as to focus on divergingradiation 109 from a radiation source. In addition the pixel structure115 within the individual detector modules 102 is angled so as to focuson the same radiation source.

The tilting of the detector modules may create unacceptable gaps betweenneighboring detector modules within the detector layer 510. These gapsare shown to be effectively filled by the complete offset of every otherdetector modules comprising the offset detector layer 510.

One implementation of a focused structure detector geometry such as aring (or partial ring) employs a front-end detector comprised of asingle layer (non-offset) or single layer with an offset layer (which isto be treated in this application as a single layer) of tilted edge-ondetector modules. As in the case of planar detectors, a focused pixelstructure can be implemented along the lengths of the edge-on tilteddetector modules (creating focused ring and focused partial ringdetector geometries).

Suitable detector configurations and materials have been previouslydescribed for Compton, PET, nuclear medicine and x-ray imaging (Nelson,et al., U.S. Pat. No. 6,583,420; Nelson, U.S. Pat. No. 7,291,841;Nelson, U.S. Pat. No. 7,635,848; Nelson, U.S. Pat. No. 8,017,906;Nelson, U.S. Pat. No. 8,115,174; Nelson, U.S. Pat. No. 8,115,175;Nelson, U.S. Pat. No. 8,183,533; Nelson, U.S. patent application Ser.No. 13/199,612, U.S. Publication No. 2012/0181437; Nelson, U.S. patentapplication Ser. No. 13/507,659, U.S. Publication No. 2013/0028379).Examples of suitable detector configurations include a single ormultilayer face-on detector, a single or multilayer edge-on detector anda multilayer detector comprised of face-on and edge-on detectors.

Edge-on detectors may incorporate SAR capability and face-on detectorsmay incorporate DOI capability. Examples of suitable detector materialsand formats previously described include semiconductor detectors,structured detectors such as structured 3D silicon (Parker S., et al.,IEEE Trans. Nucl. Sci., 53 (2006) 1676-1688; Da Via C., et al., Nucl.Instr. Meth. A594 (2008) 7) as well as other structured 3D semiconductormaterials (Diamond, Ge, Se, GaAs, CdTe, CZT, etc.), structured quantumdots (Urdaneta, M. et al., IEEE Nuclear Science Symposium, oralpresentation, 2010) and structured scintillators, and scintillators.

Structured quantum dot detectors offer flexibility since a variety ofcell shapes (including trenches) can be implemented (Nelson, U.S. patentapplication Ser. No. 13/507,659, U.S. Publication No. 2013/0028379).Furthermore, the selection of quantum dot materials can be varied as afunction of position within the substrate in order to enhance a type ofinteraction such as Compton scattering or the photoelectric effect.

Additional detector options include structured, gas-filled strawdetectors with appropriate low-Z or moderate-Z material annuli whichprovide suitable spatial and temporal resolution (Nelson, U.S. Pat. No.8,017,906), liquefied gas based detectors (such as Xenon),semiconductor-based or gas-based Medipix detectors and low temperature(such as GE and superconductor) detectors.

Multiple Compton-PET (one-sided PET) views of a volume of an object tobe imaged can be acquired as a result of detector system rotation aboutthe object to be imaged. An alternative imaging format is to rotate theobject and keep the detector system stationary. Additional objectvolumes can be imaged, if needed, by translating (typically) the objectthrough the scanner system.

It should be noted that if the Compton camera image quality isn'tsuitable for the nuclear medicine imaging applications of interest thena collimator can be inserted in front of the detector so that the systemof collimator and detector can function as a conventional SPECT/gammacamera. Since the collimator imposes a degree of directionality then theSPECT/gamma camera implementation of a Compton camera can utilize bothCompton scatter interactions (and tracking capabilities) as well asdirect photo-electric interactions (which have a much higher probabilityof occurring at lower energies such as 140.5 keV versus 511 keV in low-Zand high-Z detectors).

The direct photo-electric interactions would not be used in conventional(no electron tracking) Compton camera imaging. Furthermore, a miniatureversion of the Compton-PET detector system can be implemented as aCompton-PET hand-held detector probe. The addition of a nuclear medicinecollimator permits the Compton-PET detector probe to function as a gammacamera hand-held detector probe.

Coincidence Compton-PET detector systems extend the implementations of anon-coincidence Compton-PET detector system previously described byincluding a second Compton-PET detector system and coincidence circuitrybetween the pair of Compton-PET detector systems. For example, employinga pair of planar or partial ring Compton-PET detector systems withcoincidence circuitry.

FIG. 4 shows a perspective view of a coincidence Compton-PET detectorsystem which is comprised of a pair of planar Compton-PET detectorsystems 1000 with communications links 103 operated in coincidence forimaging an object 111 (for example, the heart). Each planar Compton-PETdetector system 1000 is positioned by an electronically controlledactuator arm 130.

For the case of a partial ring Compton-PET detector system, if asufficient number of pairs of partial ring Compton-PET detector systemsand coincidence circuitry (linking all detectors) are employed, then acomplete ring coincidence Compton-PET detector system can beimplemented. The complete ring geometry can be achieved with a singlepair of partial ring Compton-PET detector systems if each partial ringcovers an angular aperture of 180 degrees.

If the Compton scatter capability of a front-end detector is not needed(for example, if only one complete Compton camera is needed for non-PETimage applications) then there is the option of employing only aPET-compatible detector for the second detector system. Additional pairsof Compton-PET and/or PET-compatible (or combinations of both) detectorswith appropriate coincidence circuitry can be combined to form anenhanced coincidence Compton-PET detector system. (Note that a dummy ornon-functional equivalent of the front-end detector can be used to makea stand-alone PET-compatible detector unit “see” a comparable radiationfield to what the back-end detector experiences in a coincidenceCompton-PET system without the cost of an active front-end detector).

The previous description of a flexible non-coincidence Compton-PETdetector system applies to the Compton-PET detector systems used in acoincidence Compton-PET detector system. Consider the case in which atleast one of the two detector system is a Compton-PET detector system.The front-end and back-end detectors offer suitable 3D spatialresolution, energy resolution and temporal resolution. Both thefront-end and back-end detectors must provide adequate temporalresolution for an expected event rate such that accurate event trackingcan be enabled both within the front-end and back-end detectors andbetween the front-end and back-end detectors since Compton scatter andphotoelectric interactions can be recorded in both front-end andback-end detectors.

As previously described for non-coincidence Compton-PET detector systemsthis combination of front-end and back-end detectors incorporates thecapability of three conventional two-layer Compton cameras and threeconventional Compton telescope cameras. The addition of coincidencedetection capability introduces additional flexibility in that eventsinvolving a single photoelectric interaction (in which no Comptonscattering occurs) can be used for coincidence detection as well asevents involving one or more Compton scatter interactions.

Since very fast coincidence timing (TOF) can be used to improvereconstruction accuracy and reduce patient dose and/or image acquisitiontime there can be a benefit from having one or both of the front-end andback-end detectors capable of very fast timing resolution. If bothfront-end and back-end detectors are involved in the detection processthen coincidence timing can be based on using at least one of thefront-end and back-end interaction timings.

Timing resolution corrections are made for the response of one or bothdetectors (depending on whether one or both of the front-end andback-end detectors are involved in detection) and gamma ray travel timesbetween interaction locations within one or both detectors and betweendetectors (Nelson, U.S. Pat. No. 8,017,906). Commercial TOF PET systemsare capable of very fast temporal resolution (on the order of or lessthan one nanosecond).

Very fast temporal response capabilities can influence the choice ofdetector materials for front-end and back-end detectors. If thefront-end detector has a reasonable probability per photon of a Comptonscatter interaction then one option is to select a front-end detectormaterial with a very fast temporal response and select a (possibly muchless expensive) back-end detector material with a much slower temporalresponse.

If a gamma ray undergoes a Compton scatter interaction in at least oneof the front-end and back-end detectors as well as additionalinteractions such that the energy of the incident particle can beestimated then photon directionality based on the appropriate Comptoncamera reconstruction algorithm (for the Compton camera designsdescribed for non-coincidence Compton-PET detector system) can becompared with photon directionality based on coincidence (line-of-sight)between the Compton-PET detector systems operating in coincidence.

The Compton-based directionality can be used to estimate the degree ofvalidity of the coincidence (line-of-sight) assumption, includingacollinearity. This capability can be used to help reject some of thephotons that undergo Raleigh and/or Compton scattering within the objectand its surroundings as well as Rayleigh scattering or difficult todetect Compton scattering within the detectors. In addition, a(combined) non-coincidence Compton-PET (one-sided PET) reconstructedimage can be compared to a coincidence PET reconstruction image.(Nelson, U.S. Pat. No. 8,017,906). Unpaired detected events (in whichcoincidence fails since only one of the two annihilation photons isdetected and is considered legitimate) by a Compton camera can stillcontribute to the Compton scatter reconstruction image.

As previously described for the case of non-coincidence Compton-PET(one-sided PET) detector systems, system cost (in some cases) may bereduced if the back-end detector 3D spatial resolution capability islowered to 2D capability while maintaining adequate energy and temporalresolution. The 2D spatial resolution of the back-end detector impliesthat it offers limited performance as a stand-alone PET detector forgamma rays that aren't Compton scattered by the front-end detector.

The back-end detector should provide good stopping power. The Comptonscattering front-end detector offers suitable 3D spatial, temporal andenergy resolution. Single and multiple Compton scattering (as well asphotoelectric) interactions can occur in the front-end detector,allowing the front-end detector to function as a Compton camera, as aPET camera, as the first layer of a multilayer Compton camera and as thefirst layer in a multilayer PET camera in which it records the initialinteraction location, energy deposition and event timing information.(Note that if the multilayer Compton camera capability is sacrificedthen the 2D spatial resolution capability of the back-end detector canbe reduced to 1D or even block detector spatial resolution, furtherreducing costs. The back-end detector primarily provides stopping poweralong with appropriate energy and temporal resolution. The front-enddetector should offer an acceptable probability of undergoing at leastone Compton scatter interaction so that an initial location ofinteraction, timing and energy deposition can be established. If TOF PETimaging is desired then the front-end detector must offer very fasttemporal resolution. The front-end detector, due to its 3D spatialresolution capability, can still track multiple scatter interactions aswell as photoelectric events. The front-end detector retains thecapabilities of a Compton camera and a PET detector. Event trackingbetween the front-end and back-end detectors is employed.)

Multiple Compton-PET or PET views of an object volume to be imaged canbe acquired as a result of detector rotation about the object. Thealternative imaging format is to rotate the object and keep the detectorsystem stationary. If the Compton camera image quality isn't suitablefor the nuclear medicine imaging applications of interest then acollimator can be inserted in front of the detector so that the systemof collimator and detector can function as a conventional SPECT/gamma.

For the coincident and non-coincident Compton-PET configurationsdescribed there are many options for detector materials based on costand performance requirements. Assuming that acceptable-to-good energyresolution is desirable, then block, 1D, 2D and 3D back-end detectorsand 2D and 3D front-end detectors can use semiconductors, structured 3Dsemiconductors, structured semiconductor quantum dots (nanoparticles),moderate-to-bright nanophosphors, organic and inorganic scintillators,gas and liquid detectors and amplified detectors. Furthermore thesedetectors can incorporate edge-on SAR or face-on DOI (positionalencoding) capabilities.

Semiconductor and gas detectors typically offer a Fano factor noticeablyless than 1.0. If stopping power is important then sufficient detectormaterial is present in order to provide good-to-excellent attenuation.

Detector response time (for example, scintillator decay time) propertiesshould be suitable for at least event tracking at expected event rates.Very fast detectors would permit the use of TOF information to beutilized in PET reconstruction algorithms. Possible scintillators withat least one of these properties include, but are not limited to: BaFl₂,LaBr₂, LaCl₂, LSO, LYSO, GSO, GdI₃, LuI₃, SrI₂, BaHfO₃, SrHfO₃, PbWO₄,LuAP, CsI:Tl, Sm, NaI:Tl, BGO, CsI:Tl, Lu.₂O₃:Eu as well as glassscintillators, liquid scintillators and various fast-to-very fastorganic scintillators. Possible semiconductor detectors with at leastone of these properties include, but are not limited to: diamond, Si,SiC, Se, Ge, GaAs, CdTe, CZT, HgI₂, PbO, PbI₂, TlBr (as well as lownoise implementations such as silicon drift detectors or those with gainsuch as Si-APDs or SiPMs or iDADs, Se-APDs, GaAsPMs and DiamondPMs);structured 3D Si and other semiconductor materials (Parker S., et al.,IEEE Trans. Nucl. Sci., 58 (2011) 404-417) and structured semiconductorquantum dots.

A number of these semiconductor detectors can be configured as fast orvery fast photodetectors and so they can be coupled with fast or veryfast organic or inorganic scintillators. Well-known detector packages (adetector material coupled to a readout ASIC) include Medipix-baseddetectors. Additional structured detectors with gain include, but arenot limited to, gas-filled straw detectors (Nelson, U.S. Pat. No.8,017,906).

In addition, the choice of detector material can be influenced by thedetector format. For example, a 10 mm thick (or greater) CdTe or CZTface-on detector (used primarily for photo-detection) for PET imagingmay offer limited temporal resolution whereas a 1 mm thick (or less)CdTe or CZT edge-on detector (used for photo-detection and/or Comptonscattering) may qualify as a fast detector (even if SAR or DOIcorrections are not implemented). From a similar perspective a 1 mm or0.5 mm (or less) thick Si or Ge edge-on detector (used for Comptonscattering or Compton scattering and photo-detection) can be employed asa very fast detector.

If SAR or DOI capabilities are implemented to estimate the interactionlocation of an event then timing corrections can be made based on thepropagation times of electrons or holes to the anode and cathode,respectively (as described in prior patents: Nelson, U.S. Pat. No.7,635,848; Nelson, U.S. Pat. No. 8,017,906). An edge-on or face-onstructured 3D semiconductor or quantum dot detector can be employed as avery fast detector since charge propagation distances can often be lessthan 50-100 microns.

The flexibility of this Compton-PET design also allows alternativechoices for the front-end detector and back-end detector based onfactors such as lower cost and non-redundancy of features (if possible)as well as spatial resolution, energy resolution, temporal resolutionand the likelihood of Compton scatter and photoelectric interactions.For example a Compton-scatter front-end detector could be employed basedon excellent timing resolution despite reduced energy resolutioncompared to a semiconductor detector. Suitable front-end detectorcandidates with at least one of these properties include low-Z ormoderate-Z, fast and very fast organic or inorganic scintillators (orscintillating fibers) with a suitable high-speed, sensitive opticalreadout detectors (such as photodiodes, APDs, semiconductorphotomultipliers such as SiPMs and GaAsPMs, electron multiplier CCDs,microchannel plates, etc.) detectors, semiconductor-based,scintillator-based or gas-based Medipix detectors and structured,gas-filled straw detectors with appropriate low-Z or moderate-Z materialannuli (including the straw material itself) which function as a sourceof Compton electrons.

Previously, the structured straw detectors incorporated only high Zannuli in order to enhance the photoelectric effect (Nelson, U.S. Pat.No. 8,017,906). The same design technique can be used with low-Z andmoderate-Z annuli in order to enhance the Compton scatter effect.Furthermore, combinations of low/moderate-Z annuli straw detectorsfollowed by high-Z annuli straw detectors (or other high-Z detectorspreviously described) can be employed.

Detectors should offer an acceptable probability of experiencing atleast one Compton scatter interaction so that an initial location ofinteraction can be established. Event tracking within and between thefront-end and back-end detectors can be employed. If the front-enddetector offers excellent temporal resolution then TOF information canbe used to improve the reconstructed image along with a reduction inpatient dose and/or image acquisition time.

If a front-end detector lacks good energy resolution it still can beeffective if the front-end and back-end detectors offer good spatialresolution and the back-end detector offers good energy resolution.Coincidence (line-of-sight) directionality can be exploited along withthe scattered photon angle in order to estimate the incident gamma rayenergy for cases of simple Compton scatter.

Once the properties of the front-end or back-end detector have beendefined then the properties of the other detector can be selected on thebasis of which properties need to be accentuated or can be allowed todiminish (such as stopping power, energy resolution, spatial resolutionand temporal resolution). The back-end detector may primarily offerstopping power and energy resolution if the front-end detector offers 3Dspatial resolution and energy resolution. Then a cost-based decision canbe made as to whether the front-end or back-end detector (or both)should provide acceptable, fast or very fast temporal resolution.

Thus a single detector implementation does not have to embody all of thecoveted PET detector properties (high stopping power, 3D spatialresolution, fast or very fast temporal resolution). For example, thecoincidence Compton-PET detector system can implement features such asTOF imaging with a range of detector options that is much greater thanwith commercial (conventional) TOF PET systems. Partial lists of organicand inorganic scintillators and semiconductor including some of theirproperties are provided in Knoll G, Radiation Detection and Measurement,4th edition, Wiley, 2010, pages 230 (table 8.2), p. 238 (Table 8.3) andp. 492 (Table 13.3), respectively.

The flexibility of using front-end and back-end detectors for PET whichcan offer different spatial, temporal and energy resolution for PETresults in different PET images based on which detectors interact withthe pair of gamma rays from an annihilation event. For example, aCompton-PET front-end detector could Compton scatter one gamma of a pairwhich is then detected by the back-end detector. Another Compton-PETfront-end detector might fail to scatter the other gamma of the pairwhich is detected by the back-end detector. Coincidence can beestablished but the timing or spatial resolution (or both) of thefront-end detector that detects one gamma may be much better than thetiming or spatial resolution of back-end detector that detects the othergamma of the pair.

The use of front-end and back-end detectors permits flexibility as towhich detector parameters to optimize (temporal, spatial, energyresolution) as well as detector material properties (density, Comptonscatter versus photoelectric interaction probability, Compton orphotoelectric electron range) for the front-end and back-end detectors.Cost-sensitive decisions can made based on detector characteristics andgeometries in terms of how they influence various PET parametersincluding energy resolution, spatial resolution, temporal resolution,sensitivity, NECR (noise equivalent count rate), true counts,incorrectly classified events, random events, characteristic radiation,Rayleigh scatter, acollinearity, etc. For example, employ 0.5 mm thick,high-resistivity pixelated or structured (3D) Silicon arranged edge-on(adequate energy resolution, improved spatial resolution, fastertiming), rather than 1.0 mm thick, detector grade Silicon arrangedface-on. Or a material with a higher Z than Silicon could be employed toincrease photoelectric interaction probability (Ge, GaAs CdTe, etc.).

One possibility is that a front-end detector alone will be adequate. Fora dual-layer (or multi-layer) detector system all detector interactioncombinations (and thus a range of PET images with different properties)need to be considered.

CT-Compton-PET Detector Systems.

The flexibility of the Compton camera design allows it to be readilyadapted for PET (and nuclear medicine) imaging. The Compton cameradesign can also be adapted for use in diagnostic x-ray imagingapplications such as CT and projection radiography (with theunderstanding that typical data rate requirements will be much higher,spatial resolution requirements may increase, and the operational energyrange for diagnostic medical CT is typically lower than for PET andnuclear medicine imaging.

Various coincidence and non-coincidence Compton-PET detector systemimplementations have been described. An extension of this dual-useconcept is to describe a multi-use CT-Compton-PET detector system design(with the understanding that nuclear medicine imaging capability canalso be implemented).

The incorporation of CT features can be explained by examining a specialcase of a Compton-PET detector system design, the CT-Compton-PETdetector system design. This is of interest because CT-PET detectorimaging systems are commercially available. However the CT and PETdetector imaging sub-systems (which use face-on detectors) arephysically distinct. This commercial configuration involves moving thepatient with respect to the typical partial ring geometry (oralternatively a cone beam geometry) CT scanner into a physicallyseparate PET scanner. These conventional CT and PET detector sub-systemsdo not share detectors or the image acquisition space.

An alternative to the existing commercial CT-PET detector imagingsystems are improved CT-PET detector systems in which the conventionalCT scanner or PET scanner (or both) are replaced with novel edge-on CTscanners and/or PET scanners (including Compton-PET detectors) describedin this application and prior patents. For example, the traditionalface-on detector CT configuration is replaced with an edge-on CTdetector system capable of performing at least one of energyintegration, PC, and PCE (Nelson, et al., U.S. Pat. No. 6,583,420;Nelson, U.S. Pat. No. 7,291,841; Nelson, U.S. Pat. No. 7,635,848;Nelson, et al., U.S. Pat. No. 8,017,906; Nelson, U.S. Pat. No.8,115,174; Nelson, U.S. Pat. No. 8,115,175; Nelson, U.S. Pat. No.8,183,533).

A fast, improved CT-PET detector system incorporates multiple x-raytubes (two, three or more) or x-ray sources (such as carbon nanotubes,scanning electron beams, etc.) to reduce image acquisition times. NovelPET detectors include, but are not limited to, 3D crossed rod, crossedfiber-rod and encoded PET detectors.

The physically separate PET or Compton-PET scanner preferably providesone or more detector features such as suitable or excellent energyresolution, 3D spatial resolution and TOF capability. If reduced PETperformance is acceptable then one or more of energy, spatial andtemporal resolution can be degraded.

PET designs described in prior patents as well as in this patentapplication can be employed with commercial face-on CT scanners tocomprise enhanced CT-PET detector systems. Physically separatecommercial PET scanners can also be used with an edge-on CT detectorsystem in another version of an enhanced dual CT-PET imaging system.

Still another version of an enhanced dual CT-PET imaging system employsphysically separate edge-on CT and PET designs described in thisapplication and prior patents. Yet another version of an enhanced CT-PETimaging system is to employ a conventional face-on detector or edge-ondetector CT scanner with a physically separate Compton-PET detectorsystem.

An alternative to commercial and enhanced dual CT-PET detector designsare CT-Compton-PET systems in which detector components and/or space areshared, representing a cost effective and compact design compared withthe benefit that the patient remains stationary and so registrationbetween CT and PET images is straightforward. Furthermore current CTimaging sub-systems in commercial dual CT-PET systems do not offer PC orPCE capabilities which are available in enhanced dual CT-PET andCT-Compton-PET detector systems.

PC or PCE capabilities can be used for dose reduction and/ormultispectral imaging. Furthermore, multispectral imaging can beimplemented with a PC detector system by implementing x-ray tube voltageswitch (currently employed with conventional dual-energy CT detectorsystems).

CT-Compton-PET detector systems designs incorporate the capabilitiespreviously described for Compton-PET detector systems. One or morelayers of detectors can be employed.

PET options include non-coincidence (one-sided) and coincidence PETimaging capabilities. The incorporation of x-ray CT capabilities mayimpose additional requirements on the design of the radiation detectorsdepending on the energy range for the application (small animal,pediatric, adult, therapy, industrial, HLS, synchrotrons) and the event(data) rates (which, for medical CT imaging, are typically much higherthan the event rates encountered in nuclear medicine imaging).

In addition collimation may be introduced into the CT detector whichwould be of a relatively fine nature. The type and amount of collimationintroduced into the CT detector configuration is preferably sufficientto at least result in a beneficial reduction in radiation cross talkbetween detector elements during CT imaging without substantiallyreducing the efficiency of the PET detector component of the imagingsystem. If external collimation is employed to reduce the intensity ofx-rays scattered by the object from reaching the CT detector and thisexternal collimation has an undesirable impact on PET imaging efficiencyor image quality then the external collimation is preferably moveable sothat it can rotate or slide out of the detector field of view (FOV)during PET imaging. X-ray scatter correction algorithms, well-known inCT imaging, can also be employed with or without collimation along withcorrections for detector effects such as induced charges innearest-neighbor detector elements, charge cloud diffusion and radiationcross talk (energetic electrons, characteristic x-rays, bremsstrahlung)between detector elements (Nelson, U.S. Pat. No. 7,291,841; Nelson, U.S.Pat. No. 8,017,906).

If the PET detector imaging is not implemented simultaneously with theCT detector imaging then an optional movable, attenuating shield (suchas, but not limited to Cu, W, Pb, a multilayer material) can be insertedduring CT imaging to protect the PET detector from unnecessary radiationdamage and then removed during PET imaging. The insertion ofconventional nuclear medicine collimation hardware such as parallel orpinhole collimators into these Compton camera designs can provideconventional nuclear medicine imaging capabilities for those cases inwhich the Compton camera does not offer adequate imaging properties.

CT detector modes of operation can include energy integration, PC orPCE. One implementation of a CT-Compton-PET detector system is to simplyoperate the back-end PET detector independently of the front-end CTdetector and accept that the CT detector acts as an attenuator andscatterer of the 511 keV PET gamma rays.

More sophisticated CT-Compton-PET detector systems will be describednext. Implementations of detector geometries include planar (and focusedplanar) configurations and focused structure configurations such asrings and partial rings (as well as focused rings and focused partialrings). Planar, ring, and partial ring detector geometries areencountered in conventional medical diagnostic x-ray CT.

CT-Compton-PET detector systems designs described herein are based onimplementations of coincidence and non-coincidence Compton-PET detectorsystems described previously with additional constraints imposed by CTimaging. As was previously discussed, x-ray fluence rates for diagnosticmedical x-ray CT are typically sufficiently high that features such asPC and PCE are easier to implement if the distribution of detectedevents during a time interval is spread out over a number of detectorchannels. Other constraints on detector selection are related toproblems such as dose-dependent pixel performance degradation (includingpolarization issues) and detector effects described previously.

This tends to limit the selection of edge-on or face-on detector to oneor more fast-to-very fast, low-to-moderate Z semiconductors with orwithout gain capability (including, but not limited to, Si, Ge, GaAs,diamond, Se, Si-APDs, SiPMs, iDADs, Se-APDs, GaAsPMs, DiamondPMs),structured 3D semiconductors, structured semiconductor nanoparticles(quantum dots) coupled to high speed readout circuitry (such as a customreadout ASIC or a Medipix chip). Other options include configurationssuch as gas-based Medipix detectors and fast-to-very fast, brightscintillators coupled to photodetectors. Other semiconductor materialsuch as CdTe or CZT may be employed if they are sufficiently thin(typically less than 1 mm) such that issues related to polarization athigh data rates can be mitigated. Their pixel performance degradationrates and detector effects must be acceptable (or can, in part, becompensated by evaluating whether any correlated charge was deposited inneighboring pixels as in the case of the Medipix detector chip).

For the case of a focused structure detector geometry such as a ring thedetector modules can form partial rings (with detectors in a singlepartial ring that have small gaps or gaps comparable in thickness to 2Dedge-on detector plane modules (with optional collimation between thedetector plane modules). If gaps are of comparable thickness to the 2Dedge-on detector plane modules then the x-ray source is preferablycollimated to match the gaps in the detector plane and the collimatorsand detector need to move along the ring by one pixel width (detectorplane width) to acquire a complete projection for reconstruction. Thiscompensating motion and matching x-ray source collimation is not neededif at least two sets of partially-offset or completely-offset detectorrings (alternate detector modules are located at two different radii)with gaps comparable to the thickness of 2D edge-on detector modules areemployed (Nelson, U.S. Pat. No. 7,291,841).

The CT edge-on detector modules employed in a focused structure ringgeometry can also be employed in a planar CT detector geometry. One ormore layers of edge-on detector modules can be configured to be parallelor tilted with respect to adjacent detector modules in order to achievea focusing effect. As with the ring geometry implementations, layers oftilted edge-on detector modules can also be partially-offset orcompletely-offset so that tilted edge-on detector modules in a lowerlayer fill gaps between edge-on detector modules in the upper layers sothat a reasonable continuous detector is emulated. As describedpreviously a focused pixel structure can be implemented along thelengths of the edge-on tilted (or parallel) detector modules. Variousconfigurations of face-on (single or multilayer) detector modules orcombinations of face-on and edge-on detector modules, as previouslydescribed, may also be employed in planar and ring detector geometries.Optionally, SAR and DOI capabilities can be incorporated into theedge-on and face-on detector modules, respectively.

If the front-end CT detector and the back-end Compton-PET (or PET)detector operate independently of each other then the CT-Compton-PETdetector system is a limited CT-Compton-PET detector system (anintegrated limited CT-Compton-PET detector system). In this case therange of front-end CT detector designs extends from planar to focusedstructure (ring and partial ring) geometries and from traditional(low-cost) energy integration detectors to PC to PCE detectors. Thefront-end CT detector attenuates a fraction of annihilation gamma raysdirected toward the back-end Compton-PET (or PET) detector.

The planar or focused structure back-end Compton-PET (or PET) detectordoes not have to occupy the same FOV as the CT detector, larger orsmaller FOVs can be implemented according to hardware constraints, costand desired acquisition times. The back-end Compton-PET (or PET)detector can be designed to operate with a 2D or 3D spatial resolution.Non-coincidence PET (one-sided PET) imaging can be implemented with alimited CT-Compton-PET system in which the back-end detector is aCompton-PET detector. For coincidence PET imaging the back-endCompton-PET (or PET) detector can provide either 2D or 3D spatialresolution capability. Coincidence PET imaging will require the additionof a second PET detector system and the appropriate coincidencecircuitry.

If the Compton-PET detector offers 3D resolution and tracking capabilitythen both coincidence and non-coincidence PET imaging can be conductedsimultaneously. Another implementation of a limited CT-Compton-PETdetector system is to position the Compton-PET (or PET) detector outsidethe FOV of the CT detector. A radiation shield may be inserted betweenthe CT detector and the Compton-PET (or PET) detector during CToperation to limit unnecessary radiation does to the Compton-PET (orPET) detector system.

For CT-Compton-PET detector systems the front-end CT detector alsoserves as front-end detector layer for a Compton-PET detector system.The readout electronics must be suitable to handle event data rates thatare on a comparable scale to the event data rates experienced by CTdetectors or the CT detector pixel geometry must be modified to reducethe effective data rate per pixel and so reduce the requirements of thereadout electronics. The front-end and back-end detector layerspreferably include appropriate internal and intra-layer event trackingcapabilities (previously described for coincidence and non-coincidenceCompton-PET systems) depending on their intended use.

For CT applications which utilize PC or PCE capabilities several edge-onpixel geometries have been described including uniform pixel sizes (1Dor 2D pixel array) and non-uniform pixel sizes (Nelson, U.S. Pat. No.7,635,848; Nelson, U.S. Pat. No. 8,115,174; Nelson, U.S. Pat. No.8,115,175; Nelson, U.S. Pat. No. 8,183,533). Issues arise as to x-raybeam hardening with depth of penetration and the benefit of imposing amore-uniform distribution of interaction rates between pixels along thex-ray beam direction (reducing readout errors and readout electronicscosts).

If the event rate is sufficiently low a uniform pixel distribution maybe adequate even if beam hardening occurs with penetration depth. If theevent rate is high (as expected in many diagnostic medical x-ray CTapplications) and PC or PCE capability is required then a static,uniform 2D pixel array may not offer a good balance in detected eventrate per pixel unless the pixel dimensions are relatively small in termsof the stopping power of the detector material.

Unfortunately, as pixel size decreases the number of pixels and readoutelectronics increases which raises the cost of the detector modules. Inaddition to detector effects pixel readout noise can increase due toleakage issues associated with small pixels. High event rates and x-raybeam hardening with penetration depth may favor the use of a non-uniformpixel size with increasing detector depth along a pixel column. Thepixel length within a column can increased with increasing depth,resulting in a non-uniform (variable) readout element pitch in order toprovide a more-balanced count rate per pixel for the readoutelectronics. Detector pixel distributions as well as the use ofcollimating septa and/or side shielding for detector modules used in CTsystems have been described previously (Nelson, U.S. Pat. No. 6,583,420;Nelson, U.S. Pat. No. 7,291,841; Nelson, U.S. Pat. No. 7,635,848;Nelson, U.S. Pat. No. 8,017,906; Nelson, U.S. Pat. No. 8,115,174;Nelson, U.S. Pat. No. 8,115,175; Nelson, U.S. Pat. No. 8,183,533).

Furthermore, the pixel size in the axial direction (the slice thickness)can be non-uniform (benefiting dose reduction). For example, a highresolution pixel size (thin slices) could be implemented near the centerof the detector in the axial direction with a lower resolution pixelsize (thicker slices) implemented on both sides of the center.Additional non-uniform pixel size distributions can be implemented basedon imaging requirements. Additional flexibility is provided when theoutputs of two or more pixels in the axial direction can be combinedelectronically in order to synthesize the desired distribution of pixelsizes in the axial direction.

A non-uniform pixel size in the axial direction can be implemented withedge-on detectors and face-on detectors. A non-uniform pixel sizedistribution can be implemented along an arc segment. The high spatialresolution detectors can (in one implementation) be positioned at themiddle of the detector arc (that images the region of interest withinthe object being scanned) with low spatial resolution detectors oneither side.

With edge-on detectors the low spatial resolution detectors can besynthesized by combining the outputs of two or more pixels with the samecoordinates as measured with respect to the edge-on detectorsthemselves. Thus, comparable pixels from adjacent edge-on detectors(even if they are offset with respect to their neighbor) are combined.Both PCE and PC readout modes can be deployed as needed according to theimaging requirements along the axial direction and along the arc (suchas the need for energy subtraction in a limited region of image).Appropriate beam collimation and filtration can be employed to match thepixel distribution in the axial direction and along the arc.

Furthermore, non-uniformity can be extended to include the detectorgeometry type (mixing of edge-on and face-on detectors). For example,high spatial resolution edge-on detectors are (in one implementation)positioned at the middle of the detector arc (that images the region ofinterest within the object being scanned) with low spatial resolutionface-on detectors on either side (potentially reducing the over-all costof the detector system).

The principles of non-uniformity in pixel size and detector geometrytype can be applied to both ring and planar detector systems. Detectorconfigurations of reduced size can be employed if region of interest CTis implemented (retaining the high spatial resolution detectors thatimage the region of interest within the object being scanned whileeliminating the low spatial resolution detectors on either side).

A focused structure, ring geometry Compton camera design (Nelson, U.S.Pat. No. 7,291,841), may not offer optimal performance as aCT-Compton-PET camera for high event (data) rate, fan beam CT diagnosticimaging. The Compton camera would preferably use edge-on detectormodules with a uniform pixel size along a column (uniform 3D spatialresolution) whereas the PC or PCE CT system would preferably use edge-ondetector modules with a variable readout element pitch along a column.

The variable readout element pitch for CT allows the readout raterequirements of the readout ASIC-based electronics to be better balancedbetween readout elements (pixels) near the entrance surface and pixelsdistant from the entrance surface which experience a reduced beamintensity. Thus the number of readout elements can be reduced noticeablyand fewer readout ASICs of a given performance level are needed comparedto a uniform pixel array with many small pixels.

If the readout ASICs electronics offer high readout data ratessufficient to handle the maximum expected CT data rates for any pixel ina uniform pixel detector which is preferred for use in a Compton cameraor Compton-PET detector then this not an insurmountable constraint. Adrawback is a likely increase in cost due to a need for more high speedreadout ASICs than would be utilized in a dedicate CT scanner withsimilar PC or PCE capabilities but a non-uniform pixel distribution withdepth. Other issues that may arise due to this CT-COMPTON-PET detectorsystem design and the increased use of high speed readout ASICs arerelated to an increase in heat generation and therefore new coolingrequirements to avoid increased detector noise and thermal expansionissues.

There is a possibility that some readout ASICs may be moved closer tothe pixels (which may result in certain readout ASICs positioned withinthe x-ray beam path and therefore altering shielding requirements). Notethat this issue of CT detectors with uniform and non-uniform pixelarrays in CT-Compton-PET detector systems affects both the focusedstructure ring (or partial ring) detector format used in fan beam CT andthe planar detector format used in cone beam CT.

One alternative is to use readout ASICs of varying performance withrespect to depth. The highest speed readout ASICs would readout thepixels close to the entrance surface whereas readout ASICs ofprogressively slower speeds (but still sufficient for both CT andCompton camera applications) could be used to readout pixels at greaterdepths. Another alternative is to enable the edge-on detector moduleelectronics to redefine the readout element pitch according to whetherthe CT-Compton-PET detector system is functioning as a PET detectorsystem or a CT detector system.

Thus a detector module can have a selectable (fixed or variable)effective pixel width along a detector row and/or an effective pixellength along a detector column in which the effective pixel width orlength is synthesized from the outputs of one or more (typically)uniformly spaced pixels. For example, a variable, effective pixel lengthcan be optimized for CT imaging based on the beam spectrum and the beamintensity. A softer x-ray beam would preferentially be attenuated closerto the detector entrance surface than a harder x-ray beam for a givendetector material (for energies away from a detector material k-edge).

For the case of a softer x-ray beam of a given intensity the balancingof event rates between successive effective pixels in a column wouldbenefit from electronically synthesizing, relatively smaller effectivepixel lengths near the entrance surface. Relatively larger effectivepixel lengths would create a better balance of event rates betweeneffective pixels in the case of a harder x-ray beam of a givenintensity.

The advantage of a synthesized readout is that it can be optimizedaccording to the energy spectrum and the desired readout rates, thusexpanding the use of a PC or PCE CT system to a broad range of beamspectrums (applications) while retaining the uniform detector pixelgeometry useful for PET (and Compton camera) imaging. Since a SPECTcamera employs collimation to define directionality of the incidentphotons either a uniform or non-uniform detector pixel geometry can beemployed (making a CT-SPECT detector system relatively straightforwardto implement with appropriate collimation in place).

If tracking of Compton-scattered photons within the SPECT camera isimplemented then a uniform detector pixel geometry is beneficial.Features such as redefining the readout element pitch (synthesizing aneffective pixel length or width) or employing readout ASICs of varyingperformance with detector depth can be implemented in dedicated CTdetector systems as well as CT-Compton-PET detector systems and CT-SPECTdetector systems. Furthermore, CT-SPECT detector systems can employ asingle detector layer or multiple detector layers.

CT-Compton-PET detector system geometries include planar and focusedplanar detector systems and focused structure detector systems such asring and partial ring (as well as focused ring and focused partial ring)detector systems. Non-coincidence and coincidence CT-Compton-PETconfigurations are described herein based on previous descriptions ofnon-coincidence and coincidence Compton-PET configurations. The CT x-raydetectors offer suitable 3D spatial resolution, energy resolution (PCEcapability) and temporal resolution to be useful for the high x-rayfluence rates encountered in medical and non-medical CT scanning as wellas for use as the front-end detector in non-coincidence and coincidenceCompton-PET detector systems. Event tracking capability is required forCT-Compton-PET systems.

Non-coincidence CT-Compton-PET detector systems combine CT imagingcapability with one-sided PET imaging capability by employing the CTx-ray detector as the front-end detector layer that would be used in anon-coincidence Compton-PET detector system in conjunction with ahigh-stopping power back-end detector. A flexible design employsfront-end and back-end detectors that offer suitable 3D spatialresolution, energy resolution and temporal resolution.

Both the front-end and back-end detector layers provide adequatetemporal resolution for an expected event rate such that accurate eventtracking can be enabled both within the front- and back-end detectorsand between the front-end and back-end detectors since Compton scatterand photoelectric interactions can be recorded in both front-end andback-end detectors. All implementations previously described forNon-coincidence Compton-PET (three two-layer Compton cameras and threeCompton telescope cameras) are applicable with the added constraints thefront-end detector should offer suitable detection efficiency for thex-ray energy spectrums that would be used in CT imaging, it should becompatible with the event rates for CT imaging and it should offer aspatial resolution with depth that is reasonably uniform when Comptonand/or PET imaging modalities are employed.

FIG. 5 shows a perspective view of a CT-Compton-PET detector system 1000in a focused structure (partial ring) geometry which includes apoint-like x-ray 109 radiation source 125 and a gamma ray 107 radiationsource 111. The front-end detector layer 510, comprised of detectormodules 102 which use 2D pixelated array radiation detectors 115 in anedge-on geometry with base 106 and communication links 103, performs thedual role as an x-ray CT detector and a front-end detector layer(detector layer 1) for a Compton-PET detector system.

The detector modules 102 are mounted in a rigid structure 110. Theback-end detector layer 520 (detector layer 2), which could be of aplanar or focused structure geometry, is not shown. For comparison, FIG.1 can be understood to show the front-end and back-end detector layers510 and 520 (detector layers 1 and 2) for a planar CT-Compton-Petdetector geometry if the front-end detector layer 510 is suitable for CTimaging.

A reduction in cost can be realized if the Compton-PET capability isimplemented only within a sub-region of the CT detector (for example, asegment of a partial ring detector geometry or a region of a planardetector geometry). In these instances segments of CT detector modulesor regions of CT detector modules that are not involved in PET imagingdo not need to implement features such as synthesizing variableeffective pixel lengths or employing readout ASICs of varyingperformance with detector depth. Multiple Compton-PET views can still beacquired as a result of detector rotation (in some applications theobject can rotate and the detector is stationary). By reducing theactive detector area the detection efficiency will be reduced andacquisition times will, in general, increase.

Alternatively, acquisition times can be typically be reduced byincreasing the PET detector FOV beyond the CT detector FOV. As describedpreviously if the Compton camera image quality isn't suitable for thenuclear medicine imaging applications of interest then a collimator canbe inserted in front of the detector so that the system of collimatorand detector can function as a conventional SPECT/gamma camera.

Coincidence CT-Compton-PET detector systems extend the implementationsof non-coincidence CT-Compton-PET detector systems with the addition ofcoincidence detection capability by introducing a second Compton-PETdetector system along with appropriate coincidence circuitry. If theCompton scatter capability of a front-end detector is not needed thenonly a PET-compatible detector is needed for the second detector system.Implementations previously described for coincidence Compton-PETdetector systems are applicable. Thus, the detector geometries shown inFIG. 1 and FIG. 5 are applicable when employed in a coincidencedetection configuration such as FIG. 4.

Again, a reduction in cost can be realized if the coincidenceCompton-PET or coincidence PET capability is implemented only within asub-region of the CT detector (for example, a segment of a partial ringor complete ring detector geometry or a region of a planar detectorgeometry) and a matching Compton-PET or a PET-compatible back-enddetector of comparable dimensions is positioned opposite that segment orregion of the CT detector. Additional cost savings may be realized ifthe second coincidence Compton-PET detector system employs a front-enddetector that offers comparable performance to the CT detector when usedas part of a Compton-PET detector system but it lacks the extremeperformance capability of a CT detector.

Acquisition times can be typically be reduced by increasing the PETdetector FOV beyond the CT detector FOV. Multiple Compton-PET or PETviews of a limited volume of the subject can be acquired as a result ofdetector rotation about the subject. In some applications the subjectcan rotate and the detector is stationary. By reducing the activedetector area the detection efficiency will be reduced and acquisitiontimes will increase. If the Compton camera image quality isn't suitablefor the nuclear medicine imaging applications of interest then acollimator can be inserted so that the detector can function as aconventional SPECT/gamma camera.

The CT-COMPTON-PET scanner assigns the CT detector to the role of afront-end detector in a Compton-PET detector system when Compton cameraor PET (or nuclear medicine) imaging is implemented. Previously animplementation of a coincidence Compton-PET detector was describedwherein the front-end detector primarily acted as a Compton scattererand the back-end detector provided stopping power, energy resolution andtemporal resolution sufficient for event tracking with respect to thefront-end detector. Options previously described for the front-enddetector include sufficiently thin planar semiconductor detectors,structured 3D semiconductor detectors, structured quantum dotsemiconductor detectors and structured low/moderate-Z straw detectors(which typically require lower data rates than the semiconductor-baseddetectors).

If the front-end detector offers an acceptable interaction probabilitywith annihilation gammas and it is fast enough to provide the requiredcoincidence timing resolution (or very fast coincidence timing if TOFPET imaging is desired) then the back-end PET detector requirements canbe simplified since its role is primarily to detect Compton-scatteredgammas from the front-end detector. If the back-end detector is neededto provide the desired coincidence resolution (including TOF resolution)then the selection of suitable detector materials may be reduced.Reduced spatial resolution would be acceptable for a back-end detector(although Compton camera reconstruction accuracy will be reduced orlost) used in coincidence PET imaging if the front-end detector providesadequate 3D Compton-scatter information. In general, for bothcoincidence and non-coincidence Compton-PET detector system, acombination of a 3D back-end detector with a 3D front-end detector couldimprove overall detection efficiency.

PET scan times can be improved by employing additional partial-ring orplanar PET or Compton-PET detector systems that operate with or areindependent of the coincidence or non-coincidence CT-Compton-PETdetector system. These systems are referred to as enhanced coincidenceor non-coincidence CT-Compton-PET detector systems. The amount ofrotation about the object to acquire a more-complete PET image can bereduced.

Another option is to implement a coincidence CT-Compton-PET detectorsystem based on a multiple (two or more) x-ray tube or x-ray sourcesystem. For example, the angular arc of a commercial, dual x-ray tube CTpartial ring detector is approximately twice that of a single x-ray tubesystem. Multiple cone beam imaging is implemented if there are two ormore x-ray tubes or x-ray sources and corresponding planar detectors. (Awell-known example of a multi-planar detector/x-ray tube CT systemdeveloped for high speed cardiac and lung CT was the Mayo DynamicSpatial Reconstructor or DSR first implemented in the late 1970s.) Notethat if interior tomography techniques can be implemented then x-rayintensities and/or areas of planar detectors (depending on theapplication) may be reduced (Yu, H. and G. Wang, Phys. Med. Biol., Vol.54(9): pp. 2791-2805, 2009).

For the case of the focused structure partial ring geometry the CTpartial ring detector (the front-end detector) used in a dual x-ray tubeconfiguration can be split into two equal CT partial ring detectorsections so that at least one CT partial ring detector section (and itsback-end detector) can be rotated through 180 degrees when coincidencePET scanning is initiated. This could be particularly beneficial forapplications such as fast scan Cardiac CT in conjunction with CardiacPET CT. Other applications that could benefit from high resolution CTand PET or SPECT (nuclear medicine) imaging capabilities of this systeminclude head imaging and small animal imaging.

Note that the back-end detector might cover only a segment of a CTpartial ring or complete ring detector (or a region for a planardetector). If coincidence CT-Compton-PET system is implemented thesecond planar or partial ring PET detector only needs to be comparablein size to the actual PET detector implemented with the first CT planaror partial ring detector.

The efficiency of a PET detector system can be improved by addingadditional front-end detectors (and corresponding back-end detectors)adjacent to or separate from the CT partial ring detector or the CTplanar detector. These front-end detectors could utilize less demandingreadout electronics and would not require features such as pixelsynthesis since they would only be used for PET imaging and not CTimaging. Note that for the various PET implementations in which anopposing PET detector would block the x-ray beam path the opposing PETdetector is either rotated out of the beam path (the x-ray tube or x-raysource may be physically retracted when not in use) or a small openingis made in the opposing PET detector to pass the collimated x-ray beam(the PET detector rotates with the x-ray tube or x-ray source).

Multiple x-ray tubes or x-ray sources (as previously described for fast,improved CT-PET detector imaging systems) can be employed with enhancedintegrated non-coincidence or enhanced coincidence CT-Compton-PETdetector imaging systems and enhanced limited integrated CT-Compton-PETdetector imaging systems. Both stationary and rotating x-raytube-detector systems can be implemented (both designs have been usedwith dedicated CT imaging systems).

Dedicated (stand-alone) CT detector imaging systems in a ring or planardetector geometry can be implemented by reducing the functionality ofthe CT-Compton-PET detector imaging systems described herein. Aspreviously detailed, detectors with fixed (or variable) uniform ornon-uniform pixels can be implemented with the requirement that thedetectors can perform efficiently at the event count rates per pixelencountered in medical CT imaging.

CT detectors include single layer and multilayer detectors comprised offace-on detectors and/or edge-on detectors including gas, scintillator,semiconductor, low temperature (such as Ge and superconductor) andstructured detectors (such as structured 3D semiconductor, structuredquantum dot and scintillator-photodetector structured detectors). Singlelayer and multilayer detector designs of Compton cameras describedpreviously and herein can be implemented in a dedicated CT detectorimaging system with PCE capability (a simplification would be a designthat provides PC capability).

Consider a single layer, edge-on detector implementation for a medicalCT imaging system in which detector planes are aligned with the Z-axisin a ring geometry. 2D Si edge-on detectors with a wafer thickness of(for example) approximately 500 um as currently implemented may bepreferred over relatively thick, expensive, face-on CdTe or CZTdetectors in terms of operational lifetime and temporal response.

Alternative edge-on detectors of comparable thickness (approximately 500microns) which can offer improvements with respect to the stopping powerand/or temporal response performance of 2D Si at reduced cost comparedto the relatively thick, face-on CdTe or CZT detectors include, but arenot limited to, 2D GaAs, 2D CdTe and 2D CZT detectors (as well as lownoise implementations, implementations with gain or low temperatureimplementations such as 2D Ge) and structured detectors (structured 3Dsemiconductor detectors such as 3D Si, 3D GaAs, 3D CdTe, 3D CZT, 3D Ge,etc. as well as structured quantum dot detectors). In addition,detectors with thickness greater than or less than 500 microns can beimplemented depending on the image resolution requirements for the CTdetector imaging system (medical diagnostic, dental, radiation therapy,industrial, Homeland security, etc.). This single layer, edge-ondetector CT imaging system can be employed as a single layer PET imagingsystem and/or a Compton camera/Nuclear Medicine imaging system.

As described previously multiple Compton-PET implementations arepossible. Furthermore, PET and Compton camera/Nuclear Medicine imagingcan be conducted simultaneously. Depending on the fraction of the ringcircumference covered by edge-on detectors additional detectors (of thesame or different design) may need to be added to increase coincidencedetection efficiency.

For the relatively small (hardware) pixel sizes employed in currentmedical CT imaging systems Si is a reasonably efficient detector for thelower x-ray energies encountered in mammography CT and pediatric CT. Foradult CT the efficiency of Si suffers, particularly for x-ray energiesabove (approximately) 40 keV. A compromise, multilayer detectorconfiguration (for example) could employ an edge-on, 2D or structured 3DSi or GaAs (or a structured quantum dot) detector as the low-Z ormoderate-Z, front-end detector with a moderate-Z or high-Z, edge-on orface-n, back-end detector. (Note that if low temperature requirementscan be met then Ge is a candidate as a moderate-Z, face-on or edge-ondetector.)

Consider the case of an edge-on, 2D Si front-end detector. It would beof reduced height compared to a single layer, edge-on, 2D Si detectorimplementation and thus less expensive. The back-end detector (edge-onor face-on) is comprised of a moderate-Z material (such as GaAs or CdTeor CZT) or high-Z material which would emphasize photoelectricinteractions with the high energy photons that penetrate the front-enddetector.

One or more types of back-end, face-on detectors can be configured as 1Ddetectors that are positioned beneath each of the 2D Si edge-ondetectors. The thicknesses of appropriate face-on detectors should notbe so great that detrimental effects such as polarization cannot bemitigated. The cost of manufacturing such 1D detectors (material yields,butting pixels, bonding to readout electronics) should be reducedrelative to 2D detectors. More than one layer of 1D, face-on detectorscan be employed and layers can consist of the same or differentmaterials.

An alternative is to position a back-end, edge-on 1D or 2D detector(including structured 3D and quantum dot detector implementations) beloweach front end, 2D Si edge-on detector. The edge-on, 1D detector isless-costly to manufacture whereas the 2D array will typically handlehigher data rates and offer better energy resolution. This dual-layer CTdesign could be used for both low energy and high energy imagingapplications. Any combination of suitable edge-on detectors including 2Ddetectors, structured 3D semiconductor detectors and structured quantumdot detectors can be employed for the front-end and back-end detectors.

Conventional structured quantum dot detectors deploy a single quantumdot material. The use of edge-on, structured quantum dot detectorscreates an opportunity to implement a more flexible detector design. Forexample, multiple quantum dot materials can also be deployed such thatlow-Z/moderate-Z quantum dot materials are positioned near the radiationentrance surface and moderate-Z/high-Z quantum dot materials arepositioned further from the radiation entrance surface (a multilayerstructured quantum dot detector). Thus, the selection of quantum dotmaterials can be optimized for different energy ranges and the countrate per pixel as a function of distance from the radiation entrancesurface can be more-balanced.

Structured quantum dot detectors (as well as structured 3D detectors and2D semiconductor detectors) can be implemented with fixed or adjustablepixels sizes which can be uniform or non-uniform. A moderate-Z or high-Zstructured quantum dot detector can also be employed in a face-onorientation as a 1D detector positioned after a (for example) low-Z, 2DSi edge-on detector. Furthermore, moderate-Z or high-Z, fast, brightscintillator-photodetector 1D array detectors (including structureddetectors), face-on or edge-on, can be employed after a (for example)low-Z, 2D Si edge-on detector (providing limited energy resolution orsimply providing photon counting capability). The photodetector is afast, sensitive 1D photodetector chosen from (but not limited to)photodiodes, APDs, SiPMs, GaAsPMs, DiamondPMs, electron multiplier CCDsand microchannel plates with a pixel structure or a dual-readoutstructure.

Scintillator-photodetector detectors can employ scintillator screens,deposited scintillator films, and cut scintillator sheets.Scintillator-photodetector structured detectors can employ structuredscintillators (such as manufactured scintillator arrays, scintillatorsthat demonstrate columnar growth and scintillators coupled to fiberarrays) as well as scintillating or minifying scintillating, focused orunfocused, fiber arrays. Well-known scintillating fiber materialsinclude, but are not limited to, phosphors, granular phosphors,nanophosphors and quantum dots. If limited energy resolution isacceptable or only photon counting is needed for CT then a moderate-Z orhigh-Z, fast, bright, scintillator-photodetector orscintillator-photodetector structured detector (wherein thephotodetector is a fast, sensitive photodetector as previouslydescribed) can be used in place of the single layer or dual-layerdetector implementations as described herein or elsewhere (see NelsonU.S. Pat. No. 4,560,882; Nelson, U.S. Pat. No. 5,258,145; Nelson, U.S.Pat. No. 8,017,906; Nelson, U.S. patent application Ser. No. 13/507,659,U.S. Publication No. 2013/0028379).

FIG. 6 shows a minifying scintillating fiber array 140 coupled to a 1Dphotodetector 141 which is incorporated into the base unit 106. Thescintillating fiber array coupled to a photodetector readout comprises astructured detector that can be deployed in place of an edge-on detectorin a CT scanner. Adjacent structured detectors such as this can bepositioned in a continuous, partially-offset or completely offsetconfiguration. It is straightforward to extend this ring detectorgeometry comprised of an array of 1D scintillator-photodetectordetectors oriented parallel to the axial direction to multiple pixelwidths along the circumference since planar or shaped entrance surfacescintillating fiber optic arrays and small, 2D high speed photodetectorarrays are available.

The use of 1D scintillator-photodetector detectors may offer advantages(at this time) since manufacturing costs are typically reduced andbutting of 1D detectors is generally easier than butting of 2Ddetectors. The same approach applies to a planar geometry concerning theuse of 1D or 2D scintillator-photodetector detectors. Although readoutelectronics such as ASICs can be attached to the 1D or 2D photodetectorsensors externally the readout electronics can alternatively beintegrated directly on the substrate of the 1D or 2D photodetectorsensors.

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While the invention is susceptible to various modifications andalternative forms, specific examples thereof have been shown by way ofexample in the drawings and are herein described in detail. It should beunderstood, however, that the invention is not to be limited to theparticular forms or methods disclosed, but to the contrary, theinvention is to cover all modifications, equivalents, and alternativesfalling within the spirit and scope of the appended claims.

1-49. (canceled)
 50. A detector system for integrated Compton-PETradiation imaging, comprising: an x-ray and gamma ray radiation detectorhaving one or more layers of detector modules configured to generatereadout signals based on interactions of x-ray photons from an x-rayradiation source and gamma ray photons from a gamma ray radiationsource, wherein at least one layer of the detector modules includes atleast one of an edge-on detector module, a structured detector module,or a 3D spatial resolution detector module; electronics configured fortracking the interactions of at least one of the x-ray photons or thegamma ray photons and analyzing the readout signals using at least oneof energy integration or photon counting or photon counting with energyresolution techniques; and an electronic communications link to acomputer configured for post-processing, storage, and display ofradiation image data generated thereby.
 51. The detector system of claim50, wherein the detector modules provide at least one of block, 1D, 2Dand 3D spatial resolution, energy resolution or temporal resolution. 52.The detector system of claim 50, wherein the detector modules comprise acombination of edge-on and face-on detectors.
 53. The detector system ofclaim 52, wherein: the edge-on detectors incorporatesub-aperture-resolution capabilities that enhance detector spatialresolution using sub-aperture resolution (SAR) readout techniques withedge-on, side-irradiation detector geometry to estimate an interactionlocation of an event in the edge-on, side irradiation detector geometry;or the face-on detectors incorporate depth-of-interaction capabilitiesto estimate the interaction location.
 54. The detector system of claim50, wherein the detector modules provide temporal resolution adapted fortime-of-flight PET coincidence imaging.
 55. The detector system of claim50, wherein: one or more of the detector modules have pixel sizes thatare non-uniform in an axial direction, wherein a high resolution pixelsize is implemented near a center of the detector system in an axialdirection with lower resolution pixel size on both sides of the center;or one or more of the detector modules have a pixel size that issynthesized, wherein outputs of two or more pixels in the axialdirection are combined electronically to synthesize a desireddistribution of pixel sizes in the axial direction.
 56. The detectorsystem of claim 50, wherein the electronics comprise readout ASICs ofvarying performance deployed as a function of detector depth along pixelcolumns in the detector modules, wherein readout ASICs of progressivelyslower speeds read out pixels at progressively greater detector depths.57. The detector system of claim 50, wherein the electronics areconfigured for: tracking gamma ray interactions within and between twoor more layers of the detector modules; and coincidence detection of thegamma ray interactions between a pair of such detector systems facingeach other and positioned on opposite sides of an object to be imaged.58. The detector system of claim 50, wherein the electronics compriseshielded readout ASICs mounted within an etched region along a bottomedge of a semiconductor detector substrate of the detector modules. 59.The detector system of claim 50, further comprising a nuclear collimatorpositioned between the detector modules and the object being imagedthereby, and wherein the detector system is configured for 3-D imaging.60. The detector system of claim 50, wherein individual modules of thedetector modules are configured for at least one of x-ray or gamma rayphoton counting, photon counting with energy resolution, energyintegrating, or x-ray or gamma ray photon interaction tracking.
 61. Thedetector system of claim 50, wherein one or more of the detector modulescomprise structured gas-filled straw detectors with low-Z or moderate-Zannuli material selected to enhance Compton scatter.
 62. The detectorsystem of claim 50, wherein one or more of the layers of detectormodules comprise edge-on detectors irradiated from a side and having asubstrate positioned edge-on to at least one of x-rays or gamma raysincident thereon.
 63. The detector system of claim 62, wherein thedetector system is configured as a Compton-PET imaging system adapted tofunction as a SPECT imaging system or SPECT camera, or implemented witha hand-held detector probe adapted to function as a gamma camera. 64.The detector system of claim 50, wherein the x-ray and gamma rayradiation detector comprises two or more layers of the detector modules,or one or more layers of multilayer detector modules.
 65. A detectorsystem for integrated Compton-PET radiation imaging, the detector systemcomprising: an x-ray and gamma ray radiation detector having one or morelayers of detector modules configured to generate readout signals basedon interactions of x-ray photons from an x-ray radiation source andgamma ray photons from a gamma ray radiation source, wherein at leastone layer of the detector modules includes at least one of an edge-ondetector module, a structured detector module, or a 3D spatialresolution detector module; electronics configured for tracking theinteractions of at least one of the x-ray photons or the gamma rayphotons and analyzing the readout signals using at least one of energyintegration or photon counting or photon counting with energy resolutiontechniques; and an electronic communications link to a computerconfigured for post-processing, storage, and display of radiation imagedata generated thereby; wherein one or more of the layers of detectormodules comprise edge-on detector modules irradiated from a side andhaving a substrate positioned edge-on to at least one of the x-rayphotons or the gamma ray photons incident thereon.
 66. The detectorsystem of claim 65, wherein: the edge-on detector modules are configuredin a focused structure geometry comprising a focused planar, focusedring, or focused partial ring detector format, or tilted with respect toadjacent edge-on detector modules to achieve a focusing effect; or theedge-on detector modules have gaps between adjacent detector modules andthe gaps are effectively filled by offset detector modules in adjacentdetector layers.
 67. The detector system of claim 65, wherein theedge-on detector modules include at least one of a structured gas-filledstraw detector, a scintillator-photodetector, a structured quantum dotradiation detector, a structured semiconductor nanoparticle radiationdetector, a structured low temperature dielectric medium detector, or astructured 3D semiconductor detectors.
 68. The detector system of claim65, wherein at least some of the edge-on detector modules have an angledpixel structure with focused geometry relative to the radiation source.69. The detector system of claim 65, wherein the detector modules havenon-uniform geometry, and wherein the edge-on detector modules arepositioned at a middle of a detector arc that images a region ofinterest within an object, with face-on detectors on either side. 70.The detector system of claim 65, wherein one or more of the edge-ondetector modules comprises a relatively lower-Z quantum dot ornanoparticle material positioned adjacent a radiation entrance surfaceand a relatively higher-Z quantum dot or nanoparticle materialpositioned farther from the radiation entrance surface than therelatively lower-Z material.
 71. An integrated CT-Compton-PET detectorimaging system comprising: one or more layers of edge-on detectormodules irradiated from a side and having a substrate positioned edge-onto at least one of x-ray photons from an x-ray radiation source or gammaray photons from a gamma ray source, incident thereon; electronicsconfigured for tracking interactions of the x-ray photons and the gammaray photons and analyzing the readout signals using at least one ofenergy integration or photon counting or photon counting with energyresolution techniques; and an electronic communications link to acomputer configured for post-processing, storage, and display ofradiation image data generated thereby; wherein the detector imagingsystem is configured as a CT detector imaging system and a Compton-PETdetector imaging system configured to function independently, with aCT-PET detector imaging system having one or more layers of the edge-ondetector modules, and the CT detector imaging system having one or morelayers of the edge-on detector modules, and one or more of the x-ray orgamma ray sources adapted to implement multispectral imaging.
 72. Theimaging system of claim 71, wherein the one or more x-ray or gamma raysources comprise multiple carbon nanotube or scanning electron beamsources adapted for at least one of reducing image acquisition time orimplementing multispectral imaging.
 73. The imaging system of claim 71,wherein the CT-Compton-PET detector imaging system is adapted tofunction as a SPECT camera.
 74. The imaging system of claim 71, whereinthe one or more layers of edge-on detector modules comprise multilayerdetector modules irradiated from the side and having the substratepositioned edge-on to at least one of the x-ray photons from the x-rayradiation source or the gamma ray photons from the gamma ray radiationsource.
 75. The imaging system of claim 74, comprising two or morelayers of the edge-on detector modules, wherein a CT-PET detectorimaging system comprises at least a first layer of the edge-on detectormodules and a CT detector imaging system comprises at least a secondlayer of the edge-on detector modules or a layer of face-on detectormodules.